Scintillation detectors have been employed in medical imaging applications for some years and are embodied in devices called “gamma cameras” and also called “Anger” cameras after the developer, H. O. Anger. Typically, these cameras use scintillation crystals of sodium iodide doped with thallium as detectors for detecting gamma rays from radiopharmaceuticals used in nuclear medicine studies. The crystals are typically round or rectangular plates with their largest dimensions from 12 to 30″ in diameter or diagonal. These sodium iodide crystal detectors have a typical thickness of ⅜″ to ½″ (0.9 cm to 1.3 cm), sufficient for highly efficient detection of gamma rays emitted from the isotope Tc-199m, which is included in many radiopharmaceuticals and emits gamma radiation with an energy of 140 keV. For other isotopes, other crystal thicknesses may be appropriate.
A typical gamma camera detector assembly 20 is diagrammatically illustrated in FIGS. 1a and 1b. Radiation quanta 21 incident on the assembly 20 via a collimator 22 interact with the crystal 24 to produce an amount of scintillation light uniquely and nearly linearly related to the amount of energy transferred to the crystal by the radiation. The amount of energy in the incident radiation quanta may be one parameter of interest to the user and, as a result, precise determination of the energy deposited in the crystal is of considerable interest. In particular, the user is interested in knowing whether the incident quanta has the full energy of the emitting isotope or whether it has somewhat less energy, since energy loss can result from scatter in the patient or the detector. Scattered radiation degrades the quality of the image. The energy information is represented by the amount of light produced by the scintillation crystal.
As is also shown in FIGS. 1a and 1b, the typical detector assembly 20 includes an optical system for collecting the light produced in the scintillation crystal 24 and a light conversion system for electronically sensing the light collected. There are two main components in the optical system. One is an optical window 26, typically glass, which may be bonded to one face of the crystal 24 with a transparent optical adhesive 27. The other is a reflector 28 which typically covers the remaining surfaces of the crystal. FIGS. 1a and 1b also show an array of light sensors 30, typically photomultiplier tubes (PMT's), which are used to sense the light signal and convert it into an electrical signal pulse uniquely and typically nearly linearly related to the light incident on the Light sensor. Usually an optical coupling material 32 such as an optical coupling grease or potting compound is used to optically couple the PMT's to the optical window.
The package for the crystal is a moisture-impermeable housing 34 covering all sides of the crystal 24 except the portion covered by the optical window 26. This housing 34 is typically made of aluminum and is typically bonded to the edges of the glass optical window 26 with moisture-resistant adhesive 35, such as an epoxy, to provide a hermitic seal in the case of hygroscopic crystals such as Thallium-doped sodium iodide. The housing serves to prevent entrance of moisture while allowing entrance of the gamma radiation of interest. Amplification and signal processing electronics 37 are used to analyze and display the electrical information.
The user's desire for accurate energy information may be best met with an optical system that directs or redirects all the original light toward and into the light sensors, so that all light produced is collected and converted faithfully into electrical signals. Loss of light without sensing reduces the energy information available. The higher the light production and light collection efficiency, the better the energy information quality, which is often expressed as the “energy resolution” or “spectrum peak full width at half maximum”.
Reflection is important for maximizing light collection and energy information. As depicted in FIG. 1a, light from a radiation interaction (event) 40 that is originally directed away from the light sensors encounters the crystal surface and some is reflected by the index of refraction mismatch between the crystal and the surrounding air. This is represented by the light path “A” in the FIG. 1a. Some of the light may also pass through the crystal surface and encounter the reflector 28, as is represented by the light path “B” in the FIG. 1a. The reflector redirects this light so that it also has the possibility of being collected by the PMT's.
There are also other, less helpful reflections in the optical system. Light reaching the boundary between the crystal 24 (index of refraction about 1.8 for sodium iodide) and the optical coupling compound 27 (index of refraction typically 1.5 or less) is also reflected by the index of refraction mismatch between the crystal and optical coupling compound 27 as is depicted by the light path “C” in FIG. 1a. These reflections direct light away from the PMT's. Multiple reflections of the favorable and unfavorable types may occur before light is collected and sensed, or finally lost. In a typical system, the outputs of the PMT's are adjusted to be normalized, so that their sum is directly representative of the total light collection and the desired energy information.
The location of the radiation interaction usually is also of interest to the user. For this reason an array of light sensors 30 is used to sense the light, rather than a single position-insensitive light sensor. With proper design the light from a radiation interaction 40 in the crystal is distributed among several PMT's. Prior determination of the distribution functions for light among the several PMT's as a function of position allows later electronic and software determination of the location of a gamma ray interaction in the crystal. What is important for position determination is the spreading of the light, meaning its distribution in relation to the light sensor network. Roughly speaking, if light were to impinge on only a single position-insensitive Light sensor, no position information could be extracted. Also, if the light spreads among too many PMT's, the signal amplitude in each will below and therefore too easily influenced by electronic noise. A preferred situation is to distribute the light among a relatively small number of PMT's. As with energy determination, precise location of an event is aided if more light is collected, but as mentioned, the distribution is also needed for position determination. The multiple reflections mentioned earlier tend to spread the light collection area for determining event position beyond the preferred range only encompassing a few PMT's.
In Anger's method, position signals are derived from a network of weighting impedances related to a Light sensor hexagonal array and yield event positions as X and Y signals uniquely related to the X and Y Cartesian coordinate location of the radiation interaction. The sum of signals from all PMT's provides the desired energy signal. Combined with a radiation collimator and suitable electronics, such a system provides images of a patient's radio-pharmaceutical uptake for nuclear medical imaging.
The optical window 26 shown in FIG. 1a heretofore has served multiple functions. First, it transmits light from the scintillator 24 to the PMT's. Second, it may provide a hermetic barrier protecting a hygroscopic sodium iodide crystal from attack by moisture. Third, it may provide mechanical support for the crystal and light sensors. Fourth, it may provide a controlled spacing between the crystal and PMT's which allows a proper spread of light among the PMT's and provides an optimization of position and energy analysis compatible with other elements of the imaging system. The adjustment of light spreading can also be accomplished with an additional thickness of transparent light pipe or light guide material inserted between and optically coupled to the optical window and PMT's.
To use the position-sensitive method developed by Anger, the crystal and optical window are optically continuous slabs essentially free of light scattering defects, so that light propagates directly or by reflection in a predictable way with repeatable division among the PMT's. Significant defects in the crystal or the glass will cast shadows or interrupt or alter the transport of light and thereby alter the quality of, or introduce image artifacts into, the resulting images. Part of the art of successfully fabricating gamma camera radiation detectors lies in the production of high quality, large, essentially monocrystalline sodium iodide crystal plates and the fabrication of similar size pieces of high optical quality glass.
Collection of substantially all light requires highly efficient reflectors. As a matter of practice, crystal surfaces and reflectors are typically chosen for diffuse reflective characteristics, rather than specular ones. Specular reflectors tend to transport light to the edges of the crystal, while diffuse reflectors tend to provide a more favorable confined, compact light distribution which retains the light in the area of interest to permit position sensing with a relatively small number of PMT's.
Gamma camera technology has evolved since Anger's developments so that the determination of the position information may be done by techniques different from an X-Y impedance network, or ladder approach. Such alternative techniques include direct signal digitization and computer processing, nonlinear signal elements, and so forth. Essentially, though, these methods still rely on the collection of light from a large, optically homogeneous crystal in which light is distributed among a number of light sensors, so that the essential feature of the large gamma camera plate is to produce an amount of light proportional to the energy absorbed and to distribute the light reproducibly among an array of light sensors or upon a position-sensitive light sensor composed of light sensor elements. Design parameters such as the optical window thickness, the diffuse nature of the reflectors, or the degree of reflection, or all of these have an impact on tuning and optimizing the performance of the gamma camera crystal assembly in relation to the full camera sensing and computation system.
The description above of Anger's method and others has concentrated on planar imaging. There is a derivative technique wherein the camera detector, or several such detectors, are moved to several positions, images acquired, and the results processed to produce tomographic images of the radiation distribution in the patient. This is referred to as “Single Photon Emission Computed Tomography”, or “SPECT” imaging. Both SPECT and planar imaging can be affected by absorption of the primary emission in the patient. Correction for this attenuation has led to methods for measurement and calculation of attenuation correction factors. Some of these methods involve the transmission of radiation through the patient from sources in known locations, a technique which sometimes results in high count rates at the detector.
Gamma camera detector pairs or arrays have also been applied to Positron Emission Tomography, “PET” imaging, a nuclear medicine technique based on coincidence detection and positioning of positron annihilation radiation without collimators. This method of imaging operates at high count rates because there are no collimators and uses 511 keV radiation, the detection of which is enhanced by use of crystals thicker than those typically found in conventional gamma cameras.
Typical existing gamma camera detector assemblies like those described above have the disadvantage that their effective use requires large, highly transparent crystals and associated glass window pieces with a high degree of optical perfection over large areas. They must have a high degree of freedom from bubbles, inclusions, or other defects which absorb light or alter its distribution.
These gamma camera detectors also have the disadvantage that all the PMT's are optically connected to the entire crystal, so that radiation events anywhere on the face of the crystal activate large areas of the sensing and electronics analysis system. This means that the analysis system must accommodate count rates produced by the entire crystal or large portions of it. Work with transmission attenuation correction and PET studies on gamma cameras have increased the need for high rate counting. This rate limitation can be overcome partially by choosing to involve fewer PMT's in each event processing step, such as is described in U.S. Pat. No. 5,576,547. While involving fewer PMT's improves rate capability, other properties may suffer. Too few PMT's leads to light transmission beyond the sensing area in use being lost for position or energy determination purposes, or both. However, an alternative technique allowing for high rates is available. It involves essentially dividing the crystal into individual pieces, or pixels, and sensing position by noting which pixel is struck. This technique has count rate advantages, but also has cost disadvantages in that many more PMT's are required to achieve the same spatial resolution as is available from a conventional gamma camera.
The present gamma cameras cannot provide position information toward their edges because beyond about the middle of the last edge PMT in the sensing array, position information from properly located additional tubes is not available. This leads to a useful field of view which is smaller than the size of the crystal and consequently effectively to an edge region which is “dead”. To partially mitigate this effect, the optical window and edge PMT's may extend beyond the edge of the crystal by a signification fraction of the PMT radius. However, such a glass extension or overhang precludes bringing the edge of the detector crystal close to the patient, an effect which also increases the “dead” area of the gamma camera head.
The gamma camera crystal and glass are typically flat, but can be shaped by bending in one or more dimensions if necessary. While such shaping is possible, the glass and crystal do not lend themselves easily to forming shapes which conform to the body and not at all to shapes which can be changed after fabrication. Annular assemblies can also be produced, but they have the disadvantage that thermal expansion mismatches between the Nal crystal and glass window tend to stress and break the optical coupling between the window and crystal as a result of changes in temperature.
Another shortcoming of the conventional gamma camera is that the light sensitive area of the PMT light sensor, the cathode, does not typically cover the entire glass surface of the optical window. For example, if the PMT's are round, there are naturally spaces between PMT's which are not covered with cathode material. The use of closely packed arrays of square or hexagonal PMT's partly corrects this problem, but while the spaces between cathodes are reduced in area, they still exist at the glass walls of the PMT's where no cathode material is present. Square and hexagonal tubes also typically have the disadvantage of higher cost because of their more complex shape and may sometimes have poorer performance due to electron collection problems arising from that shape.
As has been seen, the need for position information requires that light be distributed among PMT's rather than being directed only to the nearest non-position sensing PMT. This distribution requirement and the detector construction details both cause light to propagate beyond the PMT's best able to determine an event's position and in some cases cause the light to escape or be absorbed, thereby being lost to the energy signal. In other words, the requirement for light distribution and position determination inherently causes light to be lost, which in turn deteriorates position determination and energy signals.
The arrangement of the camera crystal with a contiguous optical window introduces four optical boundaries into the optical system, one between the crystal and coupling compound, another between the compound to the window, a third between the optical window to another layer of coupling compound, and finally in the joining of the final coupling compound layer to the PMT. If there is any mismatch in the index of refraction of these layers, light is reflected and tends to be less useful for position determination or may be lost entirely.
From the above discussion, it can be concluded that there is a need for an improved detector which reduces or eliminates one or more of the undesirable effects present in existing designs.